Chapter 5. Biomaterials and Tissue Engineering
5.1 Scaffolds & Surfaces (S. Partap, F. Lyons, F.J. O’Brien )
I.
Introduction
a.
Introduction to tissue engineering
b.
Why are scaffolds required? 2D v 3D culture
II.
Properties of scaffolds for tissue engineering
a.
Biocompatibility
b.
Biodegradability
c.
Mechanical Properties
d.
Scaffold Architecture
e.
Manufacturing technology
III.
Biomaterials in tissue engineering
a.
Ceramics
b.
Synthetic Polymers
c.
Natural Polymers
d.
Composites
e.
Case study: Collagen scaffolds for bone tissue engineering
IV.
Scaffolds: State of the art and future directions
Address Correspondence and Reprints Requests to:
Fergal J. O'Brien, PhD
Department of Anatomy
Royal College of Surgeons in Ireland
123 St. Stephen’s Green
Dublin 2
Ireland
Phone: +353-(0)1-402-2149
Fax:
+353-(0)1-402-2355
Email: fjobrien@rcsi.ie
I. Introduction
a. Introduction to tissue engineering
Every day thousands of clinical procedures are performed to replace or repair tissues in the
human body that have been damaged through disease or trauma. Current therapies are
focused on the replacement of the damaged tissue by using donor graft tissues (autografts,
allografts or xenografts). Problems associated with this approach include shortage of
donors or donor sites, the volume of donor tissue that can be safely harvested, donor site
pain and morbidity, the possibility of harmful immune responses, transmission of disease
and rejection of grafts [1]. Alternatively, the field of tissue engineering (a phrase that is
interchangeably used with regenerative medicine) aims to regenerate damaged tissues
instead of replacing them (with grafts) by developing biological substitutes that restore,
maintain or improve tissue function [2, 3].
In native tissues, cells are held within an extracellular matrix (ECM) which guides
development and directs regeneration of the tissue, serves to organise cells in space and
provides them with environmental signals to direct cellular behaviour. The goal of tissue
engineering is to synthesise substitutes that mimic the natural ECM to help guide the
growth of new functional tissue in vitro or in vivo. At a simplistic level, biological tissues
consist of cells, signalling mechanisms and extracellular matrix.
Tissue engineering
technologies are based on this biological triad and involve the successful interaction
between three components: (1) the scaffold that holds the cells together to create the
tissues physical form, (2) the cells that create the tissue and, (3) the biological signalling
molecules (such as growth factors) that direct the cells to express the desired tissue
phenotype (Fig. 1). Tissue engineering is a multidisciplinary field that harnesses expertise
and knowledge from a variety of fields, including those of the medical profession, materials
scientists, engineers, chemists and biologists.
Fig. 1 [Insert]
b. Why are scaffolds required? 2D v 3D culture
There are differences in cell behavior in three dimensional (3-D) vs. two dimensional (2-D)
environments. In vitro 3-D cell culture conditions more accurately model in vivo biological
responses, as the conditions more closely resemble the natural structure and function of
tissues in vivo [4]. These conditions can be created by using a 3-D scaffold that acts as a
template, allowing cells to produce and deposit extracellular matrix (ECM) that would not
be possible in 2-D environments. 2-D cell culture does not allow cells to move or assemble
with the freedom they have in vivo, and thus cannot replicate the effects of nutrient
gradients, signal propagation or the development of bulk mechanical properties. Studying
these cells in 3-D models allows us to better understand their biochemical and biophysical
signaling responses as they would normally occur in vivo, particularly the external signals
occurring in the ECM, as well as the mechanical and chemical signals arising from both
adjacent and even distant cells [5]. This approach can lead to the generation of more
accurate cell- based assays for engineering of suitable biomaterials that can be used to
determine the cell-material interaction.
II. Properties of scaffolds for tissue engineering
All scaffolds for tissue engineering applications are designed to perform the following
functions: (1) to encourage cell-material interactions i.e. cell attachment, differentiation and
proliferation, eventually leading to the deposition of extracellular matrix, (2) to permit the
transport of nutrients, wastes and biological signalling factors to allow for cell survival, (3)
to biodegrade at a controllable rate which approximates the rate of natural tissue
regeneration, and (4) to provoke a minimal immune and/or inflammatory response in vivo.
The following parameters must be considered when designing a scaffold for tissue
engineering.
a. Biocompatibility
The implantation of a scaffold may elicit different tissue responses depending on the
composition of the scaffold. If the scaffold is non-toxic and degradable, new tissue will
eventually replace it; if it is non-toxic and biologically active then the scaffold will
integrate with the surrounding tissue. However, if the scaffold is biologically inactive, it
may be encapsulated by a fibrous capsule, and in the worst case scenario if the scaffold is
toxic, rejection of the scaffold and localised death of the surrounding tissue can occur [6].
Biocompatibility is the ability of the scaffold to perform in a specific application without
eliciting a harmful immune or inflammatory reaction. For a scaffold to positively interact
with cells and with minimal disruption to the surrounding tissue, it should have an
appropriate surface chemistry to allow for cellular attachment, differentiation and
proliferation. Cells primarily interact with scaffolds via chemical groups on the material
surface or topographical features. Topographical features include surface roughness and
pores where cell attachment is favoured. Alternatively, cells may recognise and
subsequently bind to the arginine-glycine-aspartic acid (RGD) cell adhesion ligand.
Scaffolds synthesised from natural extracellular materials (e.g. collagen) already possess
this specific ligands, whereas scaffolds made from synthetic materials may be designed to
deliberately incorporate them.
b. Biodegradability
The severity of an immune or inflammatory reaction is not only determined by the actual
scaffold itself, but is also dependent on the scaffold’s degradation products. Ideally,
scaffolds are designed to be completely replaced by the regenerated extracellular matrix by
integrating with the surrounding tissue, eliminating the need for further surgery to remove it
[7]. Scaffolds should degrade with a controllable degradation rate, (approximating the rate
of natural tissue regeneration), as well as with controllable degradation products. As it
degrades, the breakdown products should be non-toxic and easily excreted from the body
via metabolic pathways or the renal filtration system [8].
c. Mechanical Properties
The scaffold provides structural integrity to the engineered tissue in the short term.
Furthermore, it provides a framework for the three dimensional (3-D) organisation of the
developing tissue as well as providing mechanical stability to support the growing tissue
during in vitro and/or in vivo growth phases [9]. The mechanical properties of the scaffold
should be designed to meet the specific requirements of the tissue to be regenerated at the
defect site. Furthermore, at the time of implantation, the scaffold should have sufficient
mechanical integrity to allow for handling by the clinician, be able to withstand the
mechanical forces imposed on it during the implantation procedure and survive under
physiological conditions. Immediately after implantation, the scaffold should provide a
minimal level of biomechanical function that should progressively improve until normal
tissue function has been restored, at which point the construct should have fully integrated
with the surrounding host tissue.
d. Scaffold Architecture
Porous structures allow for optimal interaction of the scaffold with cells. The pore
architecture is characterised by pore size and shape, pore interconnectivity/tortuosity,
degree of porosity and surface area. The microstructure determines cell interactions with
the scaffold, as well as molecular transport (movement of nutrients, wastes and biological
chemicals e.g. growth factors) within the scaffold. Specifically, pore size determines the
cell seeding efficiency into the scaffold [10]; very small pores prevent the cells from
penetrating the scaffold, whilst very large pores prevent cell attachment due to a reduced
area and therefore, available ligand density. Subsequently, cell migration within a scaffold
is determined by degree of porosity and pore interconnectivity/tortuosity. A scaffold with
an open and interconnected pore network, and a high degree of porosity (>90 %) is ideal for
the scaffold to interact and integrate with the host tissue [11].
e. Manufacturing technology
In order for a scaffold or engineered construct to become commercially available in a
clinical setting, the cost effectiveness of it should be considered; particularly when it is to
be scaled up from making one at a time in a research laboratory to a production process
allowing small batch quantities of 100 to 1000 constructs to be made. In addition, as
clinicians ideally would prefer “off the shelf” products that may be used routinely, it is
important to take into consideration how the constructs will be transported and stored in
clinical environments. The cost effectiveness will be determined by the choice of
biomaterial, which will in turn affect the selection of fabrication method. Many different
techniques have been used to fabricate scaffolds for tissue engineering. The following
summarises the most commonly used methods.
Particulate Leaching Methods
Particulate leaching is a technique that uses solid particles of a particular size to act as a
template for the pores; water soluble particles are frequently used as they can easily be
leached out of the final product by simply washing the final product with water. In solvent
casting-particulate leaching, a polymer dissolved in a solvent is mixed with salt particles in
a mould; the solvent is then evaporated to give a polymer monolith embedded with the salt
particles, these are then removed by washing the scaffold with water, resulting in the
formation of a porous scaffold [12]. Another variation of this technique is melt mouldingparticulate leaching: in this particular technique the polymer is cast into a mould with the
embedded solid porogen. The polymer is set by applying heat and pressure, and again the
porogen is leached away by washing the resulting product with water to yield a porous
polymer scaffold [13].
Phase Separation
Various forms of phase separation techniques enable the creation of porous structures. A
two phase polymer system that is homogenous can become thermodynamically unstable by
altering the temperature leading to (1) liquid/liquid or (2) liquid/solid phase separations. In
the first, a polymer is dissolved in a molten solvent, a liquid/liquid phase separation (where
one phase is concentrated in polymer whilst the other is not) is achieved by lowering the
temperature. The two phase liquid is quenched to yield a two phase solid, and the solvent is
then removed yielding a porous polymer, this is known as thermally induced phase
separation (TIPS) [14]. In the second, a polymer is dispersed in a solvent which is then
frozen to induce crystallisation of the solvent to form solvent crystals that act as templates
for the pores. These crystals are then removed by freeze drying to yield a porous foam.
Manipulation of the processing conditions enables the creation of different pore sizes and
distributions [15].
Foaming
Foaming techniques use gaseous porogens that are produced by chemical reactions during
polymerisation, or are generated by the escape of gases during a temperature increase or
drop in pressure. Nam et al. 2000 [16] synthesised poly (lactic acid) [PLA] scaffolds using
ammonium bicarbonate which acted as both a gas foaming agent and as a solid salt
porogen, an increase in temperature caused the formation of carbon dioxide and ammonia
to create a highly porous foam. Also, high pressure carbon dioxide can be used to foam
polymers by saturating a prefabricated polymer monolith. A subsequent reduction in
pressure causes a decrease in solubility of the carbon dioxide within the polymer, and as the
carbon dioxide gas tries to escape it causes the nucleation and growth of bubbles resulting
in a porous microstructure [17].
Emulsion Templating
Porous structures can also be obtained by using emulsion templating techniques. The
internal phase of the emulsion acts as a template for the pores whilst polymerisation occurs
in the continuous phase (in which the monomer is dissolved). After polymerisation, the
internal phase is removed to give a templated porous material. The resulting porous
microstructures are replicas of the internal phase droplets around which polymerisations
were performed. The size of the emulsion droplets is preserved, producing polymer foams
with approximately the same size and size distributions as that of the emulsion at the point
of polymerisation [18].
Solid Free Form (SFF) Fabrication
Solid free form (SFF) fabrication or rapid prototyping (RP) technologies uses layer
manufacturing techniques to create three dimensional scaffolds directly from computer
generated files. There are a few techniques that come under this group including
stereolithography, selective laser sintering, fused depositional modeling and three
dimensional printing. However, all the techniques share the same principle where powders
or liquids are solidified one layer at a time to gradually build a three-dimensional scaffold.
The layering is controlled by computer assisted design (CAD) programs where the scaffold
architecture is designed and modelled. Data collected from computed tomography (CT) or
magnetic resonance imaging (MRI) scans may also be used to create CAD models that are
specific to the tissue to be regenerated [19].
Combination of Techniques
The techniques discussed above can also be combined with each other depending on the
exact requirements of the scaffold, e.g. phase separation (freeze drying) techniques can be
combined with emulsion templating processes. Whang et al. 1995 created an emulsion that
was quenched using liquid nitrogen, which was then freeze dried to produce porous PLGA
polymeric monoliths [20].
Fig. 2 [Insert]
III. Biomaterials in tissue engineering
A number of different categories of biomaterials are commonly used as scaffolds for tissue
engineering.
a. Ceramics
Ceramics (inorganic, non metallic materials) used within the biomedical field are classified
as being either bioinert or bioactive. The bioinert ceramics include materials such as
alumina and zirconia that are typically used as implants for musculoskeletal, oral and
maxillofacial applications whilst the bioactive group include the calcium phosphates, the
bioglasses and glass-ceramics [6]. All bioceramics are also further defined as being
osteoconductive (supporting bone growth) or osteoinductive (stimulating bone growth); all
types of bioceramics are osteoconductive as all support the formation of bone, but not all
are osteoinductive. The calcium phosphate based bioceramics, bioglasses and glassceramics are commonly used as scaffolds for bone tissue engineering as they have a
compositional similarity to the mineral phase of bone [21]. Hydroxyapatite (HA) and tricalcium phosphate (TCP) are two of the most commonly used calcium phosphate
bioceramics in tissue engineering applications. TCP is used as a degradable scaffold, whilst
HA, which is non-resorbable and has the added advantage of being osteoinductive, is
typically used for coating biomedical implants to induce bone regeneration, allowing the
implant to integrate with the surrounding tissue. For this reason, HA has shown much
popularity for use as a scaffold for tissue engineering.
b. Synthetic Polymers
The mechanical, physical and biological properties of synthetic polymers can be tailored to
give a wide range of controllable properties that are more predictable than materials
obtained from natural sources. The advantage of using synthetic materials is that the
resulting properties can be customised by adjusting the ratios of the monomer units (basic
building blocks of the final polymer) and by the incorporation of specific groups (e.g. RGD
peptide that cells can recognise). Also, the degradation rate and products can be controlled
by the appropriate selection of the segments to form breakdown products that can either be
metabolised into harmless products or can be excreted via the renal filtration system [8].
Among the many biodegradable synthetic polymers used for tissue engineering
applications, there are numerous reports on the use of polylactic acid (PLA), polyglycolic
acid (PGA) and their copolymers poly (DL-lactic-co-glycolic acid) (PLGA), which are
approved by the US Food and Drug Administration (FDA). These polymers degrade by
hydrolytic mechanisms and are commonly used because their degradation products can be
removed from the body as carbon dioxide and water. However, a disadvantage is that there
is a lowering of the pH in the localised region resulting in inflammatory responses when
they do degrade. Polycaprolactone (PCL) has a very similar structure to PLA and PGA and
is also degraded via hydrolytic mechanisms under physiologic conditions (Fig. 2). In
addition, it is degraded enzymatically and the resulting low molecular weight fragments are
reportedly taken up by macrophages and degraded intracellularly. It is predominantly used
for drug delivery devices because it has a slower degradation rate than PGA and PLA.
However, more recently, it is increasingly finding applications in tissue engineering [22].
Traditionally, polyurethanes were used in the biomedical field as blood contacting materials
for cardiovascular devices, and were intended to be used as non-degradable coatings. More
recently they have been designed to be biodegradable by being combined with degradable
polymers such as PLA for soft tissue engineering applications [14]. Poly(ethyleneglycol)
[PEG] is a biocompatible, non-toxic, water soluble polymer that is a liquid at cold
temperatures and elastic gel at 37 oC [23]. PEG based copolymers have been used as
injectable scaffolds for bone as well as for drug delivery applications [24].
Also,
copolymers of PEG and PLA have been created where the degradation rate and
hydrophilicity could be controlled by adjusting the ratio of the hydrophilic (PEG) to
hydrophobic (PLA) blocks.
Fig. 3 [Insert]
c. Natural Polymers
Natural polymers offer an alternative to synthetic polymer systems (which intrinsically lack
cell recognition signals) as they can more closely mimic the natural extracellular matrix of
tissues. Alginate and chitosan are two natural polysaccharides that do not exist within the
human body but have been investigated for tissue engineering applications because they are
structurally similar to the glycosaminoglycans (GAGs) found in the natural extracellular
matrix of tissues i.e. skin, bone, and blood vessels. Alginate originates from seaweed and is
attractive because of its low toxicity, water solubility and its simple gelation chemistry with
calcium ions. Alginate hydrogels have been investigated for use as scaffolds for cartilage
[25] and liver regeneration [26], as well as for wound dressings [27]. Chitosan is a
derivative of naturally occurring chitin which is found in the exoskeletons of crustaceans. It
has a low toxicity and is biocompatible. Chitosan scaffolds have been investigated for skin
and bone tissue engineering [28].
Given the importance of GAGs in stimulating normal tissue growth, the use of GAGs as
components of a scaffold for tissue engineering appears to be a logical approach for
scaffold development. Hyaluronic acid (sometimes referred to as hyaluronan) is one of the
largest GAG components found in the natural extracellular matrix of all soft tissues and
synovial fluid of joints [29]. The applications of pure hyaluronic acid in tissue engineering
applications are limited because of its easy dissolution in water and fast biodegradation in
biological environments. However, it can be chemically modified to produce a more
hydrophobic molecule, thus reducing its solubility in water. Hyaluronic acid scaffolds are
known to be biocompatible, and cells easily adhere to and proliferate on this material.
Hyaluronic acid also plays a significant role in wound healing and can be modified for drug
delivery applications.
Structural proteins such as fibrin are also utilised in tissue engineering applications. Fibrin
can be used as a natural wound healing material, and has found applications as a sealant
and adhesive in surgery. It can be produced from the patient’s own blood, to be used as an
autologous scaffold. However, the stability of the material is limited as it can be easily
degraded unless apronitin, a protein inhibitor, is used to control the rate of degradation.
Fibrin hydrogels have been used to engineer tissues with smooth muscle cells [30] and
chondrocytes [31]. Alternatively, gelatin (a derivative of collagen) that is produced by
altering the helical structure of the collagen molecule by breaking it into single strand
molecules) has been investigated for cartilage tissue regeneration. [32]. However, as one of
the main disadvantages of gelatin is its poor mechanical strength, it has also been crosslinked with hyaluronic acid for skin tissue engineering, and with alginate for wound healing
applications [33]. Instead, collagen, the main component found in the extracellular matrix
of mammalian connective tissues has found use in tissue engineering applications including
skin substitutes [34], scaffolds for bone and cartilage, vascular applications and as drug
delivery systems. As is typical of all natural polymers, collagen gels also display poor
mechanical properties. However, these can be improved by employing both chemical and
physical crosslinking methods. Physical crosslinking methods include UV radiation and
dehydrothermal treatments, whilst cross-linking agents such as glutaraldehyde and
carbodiimides (EDAC) can be used to produce chemically cross-linked collagen hydrogels
with improved physical properties.
Fig. 4 [Insert]
d. Composites
Due to some of the problems associated with using scaffolds synthesised from a single
phase biomaterial (eg. poor mechanical properties and biocompatibility of natural and
synthetic polymers respectively, and poor degradability of bioceramics), a number of
researchers have developed composite scaffolds comprising of two or more phases to
combine the advantageous properties of each phase. For example, polymer/ceramic
composites of poly(lactic-co-glycolic acid (PLGA) and hydroxyapatite have been
investigated for tissue engineering applications [35], whilst Cui et al. [36] have produced
tri-phasic scaffolds by depositing nano-hydroxyapatite particles onto cross-linked collagenchitosan matrices. However, even though composite scaffolds such as these have shown
some promise as grafts for bone and cartilage, each one consists of at least one phase which
is not found naturally in the body and therefore has problems with either biocompatibility
or biodegradability or both.
Table 1 summarises the different types of biomaterials
described above and lists the advantages and disadvantages of each type for use as scaffolds
in tissue engineering applications.
Table 1 [Insert]
e. Case study: Collagen scaffolds for bone tissue engineering
From an engineering viewpoint, bone is a composite material made up of both organic and
inorganic phases embedded with bone cells and blood vessels. The main components of the
organic and inorganic phases are collagen and hydroxyapatite, respectively. The collagen
fibres impart tensile strength to the bone whilst the HA crystals contribute to its stiffness.
Based on this, collagen scaffolds are currently being investigated for bone tissue
engineering applications. In our laboratory, we are currently using porous collagenglycosaminoglycan (CG) composite scaffolds which are produced using a lyophilisation
(freeze drying) process. The final pore microstructure of the scaffolds can be varied by
controlling the rate and temperature of freezing during fabrication and the volume fraction
of the precipitate [15]. We have shown that by varying the final freezing temperature
during the lyophilisation process a homologous series of scaffolds with a constant
composition and solid volume fraction with distinctly different pore sizes can be produced
[10]. Additionally, experiments performed in our laboratory using osteoblasts demonstrated
that the fraction of cells attaching to the scaffold decreased with increasing mean pore
diameter, indicating that scaffold ligand density is affected by pore size where an increase
in ligand density causes increased cell attachment. In another study, we have shown that
collagen-based scaffolds seeded with rat mesenchymal stem cells promoted differentiation
along osteogenic and chondrogenic lineages demonstrating their potential for orthopaedic
applications [37]. There is also evidence to suggest that non-seeded collagen scaffolds with
incorporated growth factors implanted into defects induce bone formation [38]. A problem
with collagen-based scaffolds, as with most natural polymer scaffolds, is their poor
mechanical properties. However, these can be improved through physical and chemical
crosslinking methods [39], and allowing bone cells to produce osteoid on the scaffolds,
enabling them to subsequently mineralise the scaffold in vitro prior to implantation, also
leads to improved mechanical properties. Alternatively, as bioceramics are mechanically
stronger and are known to enhance osteoblast differentiation and proliferation, they have
been combined with collagen scaffolds to form mineralised collagen scaffolds that support
cell growth [40, 41]. There are also reports of triphasic scaffolds made from collagen, a
bioceramic and a synthetic polymer. Scaffolds made from nano-HA, collagen and PLA
were placed in defects of rabbit radius and they integrated with the defect site within 12
weeks [42]. These studies indicate that by finding an adequate balance between pore
structure, mechanical properties and biocompatibility, a collagen-based construct can
potentially support bone growth and may have real potential for bone tissue engineering.
IV. Scaffolds: State of the art and future directions
Economic activity within the tissue engineering sector has grown five-fold in the past 5
years. In 2007, approximately 50 companies offered commercially available tissueregenerative products or services, with annual sales recorded in excess of $1.3 billion,
whilst 110 development-stage companies with over 55 products in FDA-level clinical trials
and other preclinical stages spent $850 million on development [43]. The tissue engineering
approach was originally conceived to address the gap between patients waiting for donors
and the amount of donors actually available. To date the highest rates of success have been
achieved in the areas of skin regeneration where tissue-engineered substitutes have been
successfully used in patients [44].
However, much research still remains to be performed in all aspects of tissue engineering
[45]. Cellular behaviour is strongly influenced by signals (biochemical and biomechanical)
from the extracellular matrix, the cells are constantly receiving cues from the extracellular
matrix about their environment and are constantly remodelling it accordingly. Therefore, an
appropriate three dimensional structure that is predominantly thought of as playing a
mechanical role is not enough to promote the growth of new tissue. It is important that the
scaffold provides adequate signals (e.g. through the use of adhesion peptides and growth
factors) to the cells, to induce and maintain them in their desired differentiation stage, for
their survival and growth [46]. Thus, equal effort should be made in developing strategies
on how to incorporate the adhesion peptides and growth factors into the scaffolds, as well
as in identifying the chemical identity of adhesion peptides and growth factors that
influence cell behaviour, along with the distributions and concentrations required for
successful outcomes. An example would be to incorporate angiogenic growth factors in
scaffolds for different types of tissue in an attempt to generate vascularised tissues. Tissue
vascularisation can be used to establish blood flow through the engineered tissues and
strategies involving the incorporation of vasculature, as well as innervation will be of great
importance [47]. Additionally, the incorporation of drugs (i.e. inflammatory inhibitors
and/or antibiotics) into scaffolds may be used to prevent any possibility of an infection after
surgery [48].
The field of biomaterials has played a crucial role in the development of tissue engineered
products. An alternative to using prefabricated scaffolds is to use a polymer system that is
injected directly into the defect site which is polymerised in situ using either heat [49]
(thermoresponsive polymers) or light [50] (photoresponsive polymers). The advantages for
the patient with this approach over current therapies are that injectable delivery systems fill
both regularly and irregularly shaped defects (“get a custom fit”), they represent a
minimally invasive procedure therefore avoiding surgery and the potential risks associated
with it, eliminate the need for donor tissue or a donor site, and waiting time for treatment is
reduced, as it can be used whenever treatment is required.
At present, there is a vast amount of research being performed on all aspects of tissue
engineering/regenerative medicine worldwide. Thus, as the field progresses, one of the key
challenges is to try to mimic the sophistication of the natural ECM more accurately in
synthetic substitutes. As more advanced biomaterials and bioreactors are developed, and as
research leads to more knowledge on the cell signaling mechanisms required to trigger the
chain of tissue development, we will undoubtedly get closer towards our goal of reducing
the number of patients waiting for donor tissues.
References
[1] R. Langer, Biomaterials in drug delivery and tissue engineering: One laboratory's
experience Acc Chem Res 33 (2000), 94-101.
[2] A. Atala, Tissue engineering and regenerative medicine: Concepts for clinical
application Rejuvenation Res 7 (2004), 15-31.
[3] L. J. Bonassar and C. A. Vacanti, Tissue engineering: The first decade and beyond J
Cell Biochem Suppl 30-31 (1998), 297-303.
[4] M. P. Lutolf and J. A. Hubbell, Synthetic biomaterials as instructive extracellular
microenvironments for morphogenesis in tissue engineering Nat Biotechnol 23 (2005), 4755.
[5] L. G. Griffith and M. A. Swartz, Capturing complex 3d tissue physiology in vitro Nat
Rev Mol Cell Biol 7 (2006), 211-24.
[6] L. L. Hench, Bioceramics J. Am. Ceram. Soc 81 (1998), 1705-28.
[7] J. E. Babensee, A. G. Mikos, J. M. Anderson and L. V. McIntire, Host response to
tissue engineered devices Adv. Drug Del. Rev. 33 (1998), 111-139.
[8] W. E. Hennink and C. F. van Nostrum, Novel crosslinking methods to design hydrogels
Adv Drug Deliv Rev 54 (2002), 13-36.
[9] D. W. Hutmacher, Scaffolds in tissue engineering bone and cartilage Biomaterials 21
(2000), 2529-2543.
[10] F. J. O'Brien, B. A. Harley, I. V. Yannas and L. J. Gibson, The effect of pore size on
cell adhesion in collagen-gag scaffolds Biomaterials 26 (2005), 433-41.
[11] T. M. Freyman, I. V. Yannas and L. J. Gibson, Cellular materials as porous scaffolds
for tissue engineering Prog. Mater Sci. 46 (2001), 273-282.
[12] L. Lu, S. J. Peter, M. D. Lyman, H. L. Lai, S. M. Leite, J. A. Tamada, S. Uyama, J. P.
Vacanti, R. Langer and A. G. Mikos, In vitro and in vivo degradation of porous poly(dllactic-co-glycolic acid) foams Biomaterials 21 (2000), 1837-45.
[13] S. H. Oh, S. G. Kang, E. S. Kim, S. H. Cho and J. H. Lee, Fabrication and
characterization of hydrophilic poly(lactic-co-glycolic acid)/poly(vinyl alcohol) blend cell
scaffolds by melt-molding particulate-leaching method Biomaterials 24 (2003), 4011-21.
[14] A. S. Rowlands, S. A. Lim, D. Martin and J. J. Cooper-White,
Polyurethane/poly(lactic-co-glycolic) acid composite scaffolds fabricated by thermally
induced phase separation Biomaterials 28 (2007), 2109-21.
[15] F. J. O'Brien, B. A. Harley, I. V. Yannas and L. Gibson, Influence of freezing rate on
pore structure in freeze-dried collagen-gag scaffolds Biomaterials 25 (2004), 1077-86.
[16] Y. S. Nam, J. J. Yoon and T. G. Park, A novel fabrication method of macroporous
biodegradable polymer scaffolds using gas foaming salt as a porogen additive J Biomed
Mater Res 53 (2000), 1-7.
[17] D. J. Mooney, D. F. Baldwin, N. P. Suh, J. P. Vacanti and R. Langer, Novel approach
to fabricate porous sponges of poly(d,l-lactic-co-glycolic acid) without the use of organic
solvents Biomaterials 17 (1996), 1417-22.
[18] S. Partap, J. A. Darr, I. U. Rehman and J. R. Jones, "Supercritical carbon dioxide in
water" Emulsion-templated synthesis of porous calcium alginate hydrogels Adv. Mater. 18
(2006), 501-504.
[19] E. Sachlos and J. T. Czernuszka, Making tissue engineering scaffolds work. Review:
The application of solid freeform fabrication technology to the production of tissue
engineering scaffolds Eur Cell Mater 5 (2003), 29-39; discussion 39-40.
[20] K. Whang, C. H. Thomas, K. E. Healy and G. Nuber, A novel method to fabricate
bioabsorbable scaffolds Polymer 36 (1995), 837-842.
[21] K. A. Hing, Bioceramic bone graft substitutes:Influence of porosity and chemistry Int.
J. Appl. Ceram. Technol., 2 (2005), 184-199.
[22] L. Savarino, N. Baldini, M. Greco, O. Capitani, S. Pinna, S. Valentini, B. Lombardo,
M. T. Esposito, L. Pastore, L. Ambrosio, S. Battista, F. Causa, S. Zeppetelli, V. Guarino
and P. A. Netti, The performance of poly-epsilon-caprolactone scaffolds in a rabbit femur
model with and without autologous stromal cells and bmp4 Biomaterials 28 (2007), 31019.
[23] P. J. Martens, S. J. Bryant and K. S. Anseth, Tailoring the degradation of hydrogels
formed from multivinyl poly(ethylene glycol) and poly(vinyl alcohol) macromers for
cartilage tissue engineering Biomacromolecules 4 (2003), 283-92.
[24] F. Chen, T. Mao, K. Tao, S. Chen, G. Ding and X. Gu, Injectable bone Br J Oral
Maxillofac Surg 41 (2003), 240-3.
[25] W. J. Marijnissen, G. J. van Osch, J. Aigner, S. W. van der Veen, A. P. Hollander, H.
L. Verwoerd-Verhoef and J. A. Verhaar, Alginate as a chondrocyte-delivery substance in
combination with a non-woven scaffold for cartilage tissue engineering Biomaterials 23
(2002), 1511-7.
[26] J. Yang, M. Goto, H. Ise, C. S. Cho and T. Akaike, Galactosylated alginate as a
scaffold for hepatocytes entrapment Biomaterials 23 (2002), 471-9.
[27] D. Bettinger, D. Gore and Y. Humphries, Evaluation of calcium alginate for skin graft
donor sites J Burn Care Rehabil 16 (1995), 59-61.
[28] C. Mao, J. J. Zhu, Y. F. Hu, Q. Q. Ma, Y. Z. Qiu, A. P. Zhu, W. B. Zhao and J. Shen,
Surface modification using photocrosslinkable chitosan for improving hemocompatibility
Colloids Surf B Biointerfaces 38 (2004), 47-53.
[29] J. L. Drury and D. J. Mooney, Hydrogels for tissue engineering: Scaffold design
variables and applications Biomaterials 24 (2003), 4337-4351.
[30] C. L. Cummings, D. Gawlitta, R. M. Nerem and J. P. Stegemann, Properties of
engineered vascular constructs made from collagen, fibrin, and collagen-fibrin mixtures
Biomaterials 25 (2004), 3699-706.
[31] C. J. Hunter, J. K. Mouw and M. E. Levenston, Dynamic compression of chondrocyteseeded fibrin gels: Effects on matrix accumulation and mechanical stiffness Osteoarthritis
Cartilage 12 (2004), 117-30.
[32] M. S. Ponticiello, R. M. Schinagl, S. Kadiyala and F. P. Barry, Gelatin-based
resorbable sponge as a carrier matrix for human mesenchymal stem cells in cartilage
regeneration therapy J Biomed Mater Res 52 (2000), 246-55.
[33] Y. S. Choi, S. R. Hong, Y. M. Lee, K. W. Song, M. H. Park and Y. S. Nam, Studies on
gelatin-containing artificial skin: Ii. Preparation and characterization of cross-linked
gelatin-hyaluronate sponge J Biomed Mater Res 48 (1999), 631-9.
[34] I. V. Yannas and J. F. Burke, Design of an artificial skin. I. Basic design principles J
Biomed Mater Res 14 (1980), 65-81.
[35] S. S. Kim, M. Sun Park, O. Jeon, C. Yong Choi and B. S. Kim, Poly(lactide-coglycolide)/hydroxyapatite composite scaffolds for bone tissue engineering Biomaterials 27
(2006), 1399-409.
[36] K. Cui, Y. Zhu, X. H. Wang, Q. L. Feng and F. Z. Cui, A porous scaffold from bonelike powder loaded in a collagen–chitosan matrix Journal of Bioactive and Compatible
Polymers 19 (2004), 17-31.
[37] E. Farrell, F. J. O'Brien, P. Doyle, J. Fischer, I. Yannas, B. A. Harley, B. O'Connell, P.
J. Prendergast and V. A. Campbell, A collagen-glycosaminoglycan scaffold supports adult
rat mesenchymal stem cell differentiation along osteogenic and chondrogenic routes Tissue
Eng 12 (2006), 459-68.
[38] M. Murata, B. Z. Huang, T. Shibata, S. Imai, N. Nagai and M. Arisue, Bone
augmentation by recombinant human bmp-2 and collagen on adult rat parietal bone Int J
Oral Maxillofac Surg 28 (1999), 232-7.
[39] M. G. Haugh, Jaasma, M.J. and O'Brien, F.J. , Effects of dehydrothermal crosslinking
on mechanical and structural properties of collagen-gag scaffolds J Biomed Mater Res Part
A (2008) Apr 22. [Epub ahead of print].
[40] C. V. Rodrigues, P. Serricella, A. B. Linhares, R. M. Guerdes, R. Borojevic, M. A.
Rossi, M. E. Duarte and M. Farina, Characterization of a bovine collagen-hydroxyapatite
composite scaffold for bone tissue engineering Biomaterials 24 (2003), 4987-97.
[41] A.A. Al-Munajjed and F. J. O'Brien, Development of a collagen calcium- phosphate
scaffold as a novel bone graft substitute Stud Health Technol Inform. 133 (2008). 11-20.
[42] S. S. Liao, F. Z. Cui, W. Zhang and Q. L. Feng, Hierarchically biomimetic bone
scaffold materials: Nano-ha/collagen/pla composite J Biomed Mater Res B Appl Biomater
69 (2004), 158-65.
[43] Lysaght M.J., Jaklenec A. and D. E., Great expectations: Private sector activity in
tissue engineering, regenerative medicine, and stem cell therapeutics Tissue Eng 14 (2008),
305-315.
[44] I. V. Yannas, E. Lee, D. P. Orgill, E. M. Skrabut and G. F. Murphy, Synthesis and
characterization of a model extracellular matrix that induces partial regeneration of adult
mammalian skin Proc Natl Acad Sci U S A 86 (1989), 933-7.
[45] J. P. Vacanti, Editorial: Tissue engineering: A 20-year personal perspective Tissue Eng
13 (2007), 231-2.
[46] C. A. Pangborn and K. A. Athanasiou, Growth factors and fibrochondrocytes in
scaffolds J Orthop Res 23 (2005), 1184-90.
[47] R. Langer, Tissue engineering: Perspectives, challenges, and future directions Tissue
Eng 13 (2007), 1-2.
[48] M. V. Risbud and M. Sittinger, Tissue engineering: Advances in in vitro cartilage
generation Trends Biotechnol 20 (2002), 351-6.
[49] L. Klouda and A. G. Mikos, Thermoresponsive hydrogels in biomedical applications
Eur J Pharm Biopharm 68 (2008), 34-45.
[50] K. T. Nguyen and J. L. West, Photopolymerizable hydrogels for tissue engineering
applications Biomaterials 23 (2002), 4307-14.
[51] M. J. Mondrinos, R. Dembzynski, L. Lu, V. K. Byrapogu, D. M. Wootton, P. I. Lelkes
and J. Zhou, Porogen-based solid freeform fabrication of polycaprolactone-calcium
phosphate scaffolds for tissue engineering Biomaterials 27 (2006), 4399-408.
List of Figures
Porosity
Mechanical properties
Biocompatibility
Degradability
SCAFFOLDS
Tissue Engineering
(Regenerative
Medicine)
CELLS
Stem cells
(embryonic or
adult)
Co-culture of
cells
SIGNALS
Chemical
(growth factors)
Electrical
Mechanical
Fig. 1 The tissue engineering triad; factors that need to be considered when designing a
suitable structure for tissue engineering applications.
(a)
(c)
(b)
(d)
Fig. 2 Scanning electron microscopy images of porous (a) collagen-GAG scaffolds made
by freeze drying15, (b) poly-L-lactide (PLLA) foams made by solvent casting-particulate
leaching12, (c) alginate scaffolds made by emulsion templating18 and (d) polycaprolactone–
calcium phosphate composites made by solid free form fabrication methods51.
O
CH3
O
CH
C
O
O
n
Poly (lactic acid)
CH2
O
C
O
n
Poly (glycolic acid)
O
O
(CH2)5
O
n
Poly (caprolactone)
CH2
CH2
n
Poly (ethylene glycol)
Fig. 3 Chemical structures of some biodegradable synthetic polymers used as scaffolds in
tissue engineering applications
O
OH
HProperties
O
O
Scaffold
O
O
O
HO
OH
Advantages
O
HO
Disadvantages
O
O
O
O
HO
NH2
n
Alginate
n
Chitosan
OH
O
HO
O
O
O
O
HO
OH
O
HO
NH
O
n
Hyaluronic Acid
Fig. 4 Chemical structures of some natural polymers used as scaffolds in tissue engineering
applications
Bioceramics
Hydroxyapatite (HA)
Found naturally as a component
of mineral phase of bone
Biocompatible
Osteoinductive
Compositional similarity to
mineral phase of bone
Biocompatible
Biodegradable
Poly(lactic acid), Poly(glycolic
acid) and their copolymers
Mechanical and degradation
properties can be tuned by
varying polymer segments
Biocompatible
Degradation
products are CO2
and H2O creating
local acidic
conditions
Poly(ethylene glycol)
Used as an injectable gel
Mechanical and degradation
properties can be tuned by
varying polymer segments
Biocompatible
Hydrophilic
Poor cell adhesion
Collagen
Component of natural
extracellular matrix (ECM)
Poor mechanical
properties
Hyaluronic acid
Plays role in natural wound
healing
Component of natural ECM
Alginate
Originates from seaweed
Structurally similar to natural
glycosaminoglycan’s (GAG)
Biocompatible
Good cell
recognition
Biocompatible
Easily
functionalized
Good cell
recognition
Biocompatible
Simple gelation
methods
Tricalcium Phosphate (TCP)
Non resorbable
Poor mechanical
properties
Poor mechanical
properties
Synthetic polymers
Natural polymers
Poor mechanical
properties
Poor mechanical
properties
Composites
Polymer - Ceramic
Natural or synthetic polymers
combined with ceramics
Often combined for bone tissue
engineering applications
Compromise
between ‘best’
qualities of
individual
components with
overall scaffold
properties
Polymer - Polymer
Combinations of (1) syntheticAbility to tailor
Compromise
synthetic, (2) synthetic – natural mechanical,
between ‘best’
and (3) natural – natural
degradation and
qualities of
polymers possible
biological
individual polymers
properties
with overall
scaffold properties
Table 1 Properties, advantages and disadvantages of biomaterials used as scaffolds in
tissue engineering applications
Ability to tailor
mechanical,
degradation and
biological
properties
28